Controlling power in a prosthesis or orthosis based on predicted walking speed or surrogate for same

ABSTRACT

In some embodiments of a prosthetic or orthotic ankle/foot, a prediction is made of what the walking speed will be during an upcoming step. When the predicted walking speed is slow, the characteristics of the apparatus are then modified so that less net-work that is performed during that step (as compared to when the predicted walking speed is fast). This may be implemented using one sensor from which the walking speed can be predicted, and a second sensor from which ankle torque can be determined. A controller receives inputs from those sensors, and controls a motor&#39;s torque so that the torque for slow walking speeds is lower than the torque for fast walking speeds. This reduces the work performed by the actuator over a gait cycle and the peak actuator power delivered during the gait cycle. In some embodiments, a series elastic element is connected in series with a motor that can drive the ankle, and at least one sensor is provided with an output from which a deflection of the series elastic element can be determined. A controller determines a desired torque based on the output, and controls the motor&#39;s torque based on the determined desired torque.

CROSS REFERENCE TO RELATED APPLICATIONS

This Application is a continuation of U.S. patent application Ser. No.13/079,564 filed Apr. 4, 2011 which claims the benefit of U.S.Provisional Applications 61/320,991 filed Apr. 5, 2010, 61/422,873 filedDec. 14, 2010, and 61/432,083 filed Jan. 12, 2011, each of which isincorporated herein by reference.

BACKGROUND

U.S. published patent applications 2010/0174384 (“the '384 application”)and 2006/0249315, each of which is incorporated herein by reference,describe that the gait cycle for walking can be divided into fivephases: controlled plantarflexion, controlled dorsiflexion (CD), poweredplantarflexion (PP), early swing, and late swing, as depicted in FIG. 1.

The '384 application also discloses a number of embodiments oflower-extremity prosthetic and orthotic systems in which the reflextorque generation during PP is achieved via non-linear, positivefeedback between the series elastic element (SEE) motor torque and ankletorque. More specifically, the reflex action involves behaving like anon-linear spring during CD and like a torque source during PP. Thisreflex action can be implemented by driving the motor using thefollowing equation:

Motor Torque=pff×(normalized_Torque)^(n)   1

Where, pff is the power control gain tuned for high walking speed;normalized_Torque is the ankle torque, FA, normalized by a torque, To,(strongly related to users' weight); n is the power exponent, typicallyin the range of between 3 and 5 for level-ground walking. Note that pffhas units of N-m, and the value of pff controls the magnitude of thelevel of the torque reflex during fast walking. Once the desired motortorque is determined, the drive current can be computed based on theequation Motor Current=Motor Torque/kt, where kt is the motor torqueconstant. While using Equation 1 does provide good results, the resultsprovided by the control approach described below are significantlybetter.

SUMMARY OF THE INVENTION

One aspect of the invention is directed to an ankle-foot prosthesis ororthosis apparatus. The apparatus includes a shank member and a footmember that is operatively configured with respect to the shank memberso as to supporting walking and permit the foot member to plantarflexand dorsiflex with respect to the shank member. A motor is configured toplantarflex the foot member with respect to the shank member, and aseries elastic element is connected between at least one of (a) themotor and the shank member and (b) the motor and the foot member, Thereis at least one first sensor having an output from which a walking speedof an upcoming step can be predicted, and at least one second sensorhaving an output from which ankle torque can be determined. Theapparatus also includes a controller configured to control the motor'storque, based on the output of the at least one first sensor and the atleast one second sensor, so that the motor's torque for slow walkingspeeds is lower than the motor's torque for fast walking speeds.

Another aspect of the invention is directed to a method of modifyingcharacteristics of an ankle-foot prosthesis or orthosis apparatus. Themethod includes the steps of predicting what a walking speed will beduring an upcoming step and modifying a characteristic of the apparatusduring the upcoming step in situations when the predicted walking speedis slow. The modification of the characteristic results in a reductionin net non-conservative work that is performed during the upcoming stepas compared to the net non-conservative work that is performed when thepredicted walking speed is fast.

Another aspect of the invention is directed to an apparatus thatincludes a proximal member and a distal member that is operativelyconnected with respect to the proximal member by a joint so that anangle between the distal member and the proximal member can vary. Amotor is configured to vary the angle between the distal member and theproximal member, and a series elastic element is connected between atleast one of (a) the motor and the proximal member and (b) the motor andthe distal member. There is a least one first sensor having an outputfrom which a walking speed of an upcoming step can be predicted, and atleast one second sensor having an output from which a joint torque canbe determined. The apparatus also includes a controller configured tocontrol the motor's torque, based on the output of the at least onefirst sensor and the at least one second sensor, so that the motor'storque for slow walking speeds is lower than the motor's torque for fastwalking speeds.

Another aspect of the invention is directed to an ankle-foot prosthesisor orthosis apparatus that includes a shank member and a foot memberthat is operatively configured with respect to the shank member so as tosupporting walking and permit the foot member to plantarflex anddorsiflex with respect to the shank member. A motor is configured toplantarflex the foot member with respect to the shank member, and aseries elastic element is connected between at least one of (a) themotor and the shank member and (b) the motor and the foot member. Theapparatus also includes at least one sensor having an output from whicha deflection of the series elastic element can be determined, and acontroller configured to determine a desired torque based on the output,and to control the motor's torque based on the determined desiredtorque.

Another aspect of the invention is directed to a method of controllingan ankle-foot prosthesis or orthosis having a foot member and shankmember, with a motor configured to plantarflex the foot member withrespect to the shank member and a series elastic clement in series withthe motor. The method includes the steps of sensing a position of themotor, determining a deflection of the series elastic element while themotor is at the position sensed in the sensing step, and controlling themotor's torque based on the motor position sensed in the sensing stepand the deflection determined in the determining step.

Another aspect of the invention is directed to an apparatus thatincludes a proximal member, a distal member that is operativelyconfigured with respect to the proximal member so that an angle betweenthe distal member and the proximal member can vary, and a motorconfigured to vary the angle between the distal member and the proximalmember. A series elastic element is connected between at least one of(a) the motor and the proximal member and (b) the motor and the distalmember, and at least one sensor having an output from which a deflectionof the series elastic element can be determined. The apparatus alsoincludes a controller configured to determine a desired torque based onthe output, and to control the motor's torque based on the determineddesired torque.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic illustration of the phases of a user's gait cyclewhen walking on level ground.

FIG. 2A depicts the statistic range of net non-conservative work vs.walking speed for healthy human ankles.

FIG. 2B depicts the statistic range of peak-power vs. walking speed forhealthy human ankles.

FIG. 2C shows the net non-conservative work vs. walking speed when twodifferent equations are used to control a motor.

FIG. 2D shows peak-power vs. walking speed when two different equationsare used to control a motor.

FIG. 3A depicts the relationship between walking speed of the upcomingstep and the shank angular rate.

FIG. 3B depicts what shank angular rate is used in FIG. 3A.

FIG. 4A depicts one suitable gain function for use in controlling themotor.

FIG. 4B depicts another suitable gain function.

FIG. 5A is a block diagram of an embodiment that relies on torquesensing.

FIG. 5B depicts a mechanical configuration for the FIG. 5A embodiment.

FIG. 6A is a block diagram of an embodiment that relies on deflectionsand torque vs. deflection characteristics.

FIG. 6B depicts mechanical configuration for the FIG. 6A embodiment.

FIG. 6C depicts a section view of the FIG. 6B configuration.

FIG. 7 depicts a test fixture for measuring torque vs. deflectioncharacteristics.

FIG. 8A is a graph from which a spring rate can be determined

FIG. 8B is a graph depicting changes in a torque component over time.

FIG. 9 depicts the torque vs. deflection characteristics for a serieselastic element.

FIG. 10 is a Γ-Θ plot for the stance-phase torque-angle response of anintact ankle.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

In healthy humans, the ankle-foot normally creates the positive net-workand peak-power on each stride that the body needs to achieve ordinarywalk with metabolic efficiency. The net-work and peak-power in the ankleduring the stance of gait is highly related to walking speed. FIGS. 2Aand 2B depict this relationship. More specifically, FIG. 2A shows thestatistic range (+1 sigma bounds) of net non-conservative work vs.walking speed, which lies between the lines 11, 12. FIG. 2B shows theestimated statistic ranges (+1 sigma bounds) of the peak-power vs.walking speed as lines 16, 17. FIG. 2B also shows the mean value ofpeak-power vs. walking speed (as measured in a study) as line 18, whichlies between lines 16 and 17.

The data points depicted by stars in FIG. 2C shows the netnon-conservative work vs. walking speed when Equation 1 above is used tocontrol the motor current. Note that net non-conservative work can bedetermined by calculating the loop area, over one cycle of ankle-torquevs. ankle angle (e.g., as seen in FIG. 10, starting at point 1, passingthrough points, 2, 3, and 4 in sequence, and returning to point 1. Itcan be seen that the net non-conservative work is higher than thestatistic range bounded by lines 11, 12 for intact ankles, and thedeviation from that range is larger at slower walking speeds than it isat faster walking speeds. Similarly, the data points depicted by starsin FIG. 2D show the peak power vs. walking speed when Equation 1 aboveis used to control the motor current. It can be seen that the peak poweris higher than the mean value line 18 for intact ankles. The net work isalso higher, and is wasted, causing extra heat and reduction in batterylife.

To more closely mimic the human ankle-foot biomechanics for ordinarywalk across a wide range of walking speeds, the embodiments disclosed inthe '384 application may be modified by using the power control approachdescribed herein so as to deliver network and peak-power on each stridethat more closely matches the statistic ranges bounded by the lines 11,12 in FIG. 2A, and the mean line 18 in FIG. 2B, In this approach, aprediction of the walking speed for the upcoming step is made, and thatpredicted walking speed is used to set the ankle control parameters(including setting of the power control gain) for the upcoming step.

One way to predict the walking speed of the upcoming step is based onthe shank (pitch) angular rate ω_(x) based on the relationship depictedin FIG. 3A. These two velocities are highly linearly correlated suchthat the peak angular rate in stance phase serves as an excellentprediction of the walk speed of the up coming step. The correlationbetween walking speed and the shank angular rate is present at varioustimes during the stance and swing phase, but it is preferable tominimize the latency between the walking speed estimate and when it willbe applied. One way to accomplish this is to sample the shank angularrate at the very start of controlled dorsiflexion (i.e., at foot-flat),immediately before the reflex begins. This reduced latency ensures thata reflex is not applied in certain situations, such as when the user isstopping. If, on the other hand, a stale walking-speed prediction wereused, (e.g., by estimated walking speed from the shank angular rate atthe prior toe-off), the estimate might be invalid (e.g., in situationswhere the user decides to stop suddenly).

The shank angular rate may be measured by any suitable means, such as aninertial measurement unit (IMU) or an angular rate sensor (ARS). The IMUor ARS may be placed onto the top part of the prosthesis or orthosisthat is rigidly connected to a socket such that shank angular rate, asdepicted in FIG. 3B, can be measured. In alternative embodiments, itcould be mounted on the foot structure. An example of a suitable angularrate sensor is the Invensense IDG-300. In one preferred embodiment, theIMU can be made from three orthogonally-aligned angular rate sensorssuch as the Analog Devices ADXRS610, and three orthogonally-alignedaccelerometers such as the Freescale MMA7360L.

An advantage of using the angular rate sensing technique is that itprovides an instantaneous measure of angular rate just prior to invokingthe reflex control. More specifically, the maximum angular rate in thestance phase can be calculated and employed to adjust the reflex torqueresponse during the controlled dorsiflexion and powered plantar flexionphases of a step. This reflex is largely responsible for generating thenet-work and peak-power that meet human ankle-foot needs for ordinarywalking.

The reflex torque generation is achieved via non-linear, positivefeedback between the series elastic element (SEE) motor torque and ankletorque by controlling the motor using the following equation:

Motor Torque=Kv(ω_(x))×pff×(normalized_Torque)^(n)   2

where Kv(ω_(x)) is a power control gain function related to the maximumangular rate, an example of which is depicted in FIG. 4A; pff is thepower control gain tuned for high walking speed; normalized_Torque isthe ankle torque, Γ_(A), normalized by a torque, Γ₀, (strongly relatedto users' weight); and n is the power exponent, typically in the rangeof between 3 and 5 for level-ground walking. This is similar to Equation1 above, except that the right side of the equation is multiplied by again function Kv(ω_(x)) that is selected to reduce the motor torque forlower angular velocities, which correspond to slower walking speeds.Note that the companion equation for converting a desired motor torqueto a drive current for the motor remains the same for all embodimentsdescribed herein (i.e., Motor Current=Motor Torque/kt, where kt is themotor torque constant).

One suitable gain function Kv(ω_(x)) is depicted in FIG. 4A, whichstarts at 0 when the angular rate is zero, and increases linearly to 1at an angular rate ω_(TH) that corresponds to a fast walking speed.Above that threshold angular rate ω_(TH), the gain function Kv(ω_(x))remains at 1. A suitable setting for the threshold ω_(TH) is an angularrate that corresponds to a fast walking speed (e.g., an angular ratethat corresponds to a walking speed of between 1.5 and 1.75 meters persecond). In some embodiments, the threshold point may be settable by aprosthetist, preferably constrained to some legal range (e.g., to anangular rate that corresponds to a walking speed of between 1.25 and 2meters per second). In other embodiments, provisions for adjusting theω_(TH) set point within a legal range may even be made available to theend user.

The result of multiplying the right side of Equation 2 by Kv(ω_(x)) isthat the motor will be driven by lower currents for slower walk speeds.That will result in less torque at slower walk speeds (as compared towhen Equation 1 is used). When this approach is used to control aprosthetic or orthotic ankle, during the flat-foot portion of the gaitthe torque will initially be zero. The ankle torque Γ_(A) will start toincrease at the end of the controlled dorsiflexion phase. In response tothe rising Γ_(A), the controller will drive the motor based on Equation2, which will increase the torque further in a positive feedback reflexresponse. This positive feedback continues until prior to toe-off as thelower leg begins to lift the foot off the ground. At this point thepositive feedback is diminishing, so the torque starts to drops off. Thepositive feedback is quenched at toe-off because at that point there isnothing to push against, which makes the torque fall off rapidly. Inaddition, the state machine that controls the application of the reflexalso transitions to the swing phase where position control is used. Notethat operation of the state machine is described in the '384application, which is incorporated herein by reference.

The speed based power control method of Equation 2 has been implementedand tested on an iWalk™ Powerfoot™ BiOM™ prosthetic ankle/foot. WhenEquation 2 was used to control the motor, the net non-conservative workvs. walking speed is depicted by the circle data points in FIG. 2C. Acomparison between the circle data points and the star data points(discussed above) in FIG. 2C reveals that the net non-conservative workis closer to the statistic range bounded by lines 11, 12 when Equation 2is used. Similarly, the circle data points in FIG. 2D show the peakpower vs. walking speed when Equation 2 above is used to control themotor current. It can be seen that the peak power when Equation 2 isused is much closer to the mean value line 18 than when Equation is used(indicated by the star data points in FIG. 2D). This experiment resultwas obtained from a patient with weight of 240 lb and shank length of 53cm. The walk speed was measured using IMU systems, and ranged from 0.8m/s to 1.5 m/s. The system provided smooth transitions of power whenusers randomly changed their walking velocities.

In alternative embodiments, gain functions with other shapes may be usedinstead of the ramp depicted in FIG. 4A. Preferably, all such functionsstart at 0 when ω_(x)=0, end at 1, and are monotonically increasing.Examples of suitable shapes for the gain function include shapes thatresemble (a) the first quadrant of a sine curve; or (b) the third andfourth quadrants of a cosine curve (scaled and offset so as to start at0 and end at 1). Other transition shapes, including smooth shapes andshapes with abrupt changes, may also be used. For example, the curvedepicted in FIG. 4B would operate to keep the power low for low walkingspeeds (which would be suitable in certain situations like a classroom),and increase it only if the speed goes over a threshold ω_(TH2).Optionally, the gain function may also be operative for negativevelocities to control the reflex response when walking or runningbackward. For this reason, negative velocities arc included in FIG. 4B.If desired, the maximum gain for negative velocities may be lower than1, so as to provide a smaller power boost when walking backwards In someembodiments, the gain function could also be made to be a function ofvelocity when side-stepping or hopping sideways.

In some embodiments, a user interface may be provided to give theprosthetist control over the value of n in Equation 2, preferablyconstrained within some legal range (e.g., between 2 and 7). Set pointsof between 3 and 5 have been found to be preferable. Sincenormalized_Torque is Γ_(A) normalized by Γ₀, when n is high (e.g.,around 5), the current will not rise until Γ_(A) gets closer to Γ₀. Thisdelays (in time) the onset of the positive feedback. Conversely, when nis lower (e.g., around 3), the current will start to increase beforeΓ_(A) gets too close to Γ₀. This advances (in time) the onset of thepositive feedback. When the system is configured to give the prosthetistcontrol over n, n can be adjusted (e.g., based on verbal feedback fromthe end user) to maximize the user's comfort. In other embodiments, auser interface may be provided to give the end user control over n(within a legal range).

In alternative embodiments, the reflex torque generation equation may bemodified to be as follows:

Motor Torque=Kv(ω_(x))×pff×(normalized_Torque)^(nf(ω) ^(x) ⁾   3

Equation 3 is very similar to Equation 2, except that in Equation 3, theexponent n of the normalized_Torque is multiplied by a function of theangular rate ω_(x). The function f(ω_(x)) is preferably selected so thatthe resulting exponent is larger at higher angular velocities than it isat lower angular velocities. This would operate to advance the onset ofreflex (in time) when the user is walking faster, with respect to thetiming when the user is walking slower.

Note that in the embodiments described above, the system does notexplicitly make a prediction of the walking speed for the upcoming step.Instead, the system relies on the angular rate ω_(x) of the shank(which, as described above, is correlated to the predicted walkingspeed). In this case, the angular rate ω_(x) of the shank serves as asurrogate for the walking speed. In alternative embodiments, instead ofrelying on the angular rate ω_(x) of the shank, other parameters may beused to predict the walking speed. The ankle power would then beadjusted accordingly based on the predicted walking speed based on thesealternative sensors. For example, the angular rate of the leg sectionabove the knee, or the knee linear moving velocity in stance phase maybe used to predict the walking speed of the upcoming step. The Cartesiantrajectory of the ankle or knee, tracked using an IMU, could also beused to predict the walking speed of the upcoming step.

In other embodiments, the equations may implemented so as to explicitlycompute the estimated walking speed as an intermediate result, and thenadjust the various parameters based on that intermediate result tocontrol the power and net non-conservative work (e.g., by replacingKv(ω_(x)) with Kv(speed) in Equation 2).

Preferably, the system includes special-event handing to modify thepower level when it determines that a special walking environmentexists. For example, the power may be increased for upstairs/up-rampwalking, even though the walk speed is low. Or the power may bedecreased for down stairs or down ramp walking even though the walkspeed is high. Note that the ankle trajectory or knee trajectory(determined, for example, using an IMU) may be used as a discriminatorto determine if a special walking environment exists, so that thecharacteristics of the ankle (including the reflex) can be adjusted forthe special walking environment.

The system described above provides users improved net-work andpeak-power to achieve normal biomechanics for ordinary walking across arange of walking speeds. The system also uses reduced motor current atlow walking speeds, which is the case for the majority of walking inmost people's routines. This may help keep the motor temperature low,save energy, and reduce the frequency of recharging batteries and theneed to carry spare batteries. Lower currents also reduce the stress andfatigue on the drive transmission, including the series-spring, and canincrease the design life of various components in the device.

The embodiments described above rely on the ankle torque Γ_(A) as aninput to the equations that ultimately control the motor current duringcontrolled dorsiflexion and powered plantar flexion. This ankle torqueΓ_(A) may be determined by a number of approaches. One such approach,which is described in the '384 application, is to actively measure theankle torque Γ_(A) using, for example, strain gauges arranged in aWheatstone bridge configuration to measure the torque applied by thesocket attachment at the top of the ankle prosthesis.

FIG. 5A is a system block diagram for this embodiment. The prosthetic ororthotic ankle/foot includes a shank member 52 and a foot member 54operatively connected to permit plantarflexion and dorsiflexion, e.g.,by a joint 53. A motor 56 is affixed to the shank member 52, and aseries elastic element 58 sits between the shank member 52 and the footmember 54, so that it will be in series with the motor, as explained inU.S. Pat. No. 5,650,704, which is incorporated herein by reference.Driving the motor in one direction or the other will plantarflex ordorsiflex the foot member 54 with respect to shank member 52. Inalternative embodiments (not shown) the positions of the motor 56 andthe series elastic element 58 could be swapped, in which case the motorwould be mounted to the foot member 54.

A torque sensor 66 measures the ankle torque Γ_(A) and send an outputthat represents that torque to the controller 68. The controller 68 isprogrammed to control the motor 56 by implementing Equation 2. Inalternative embodiments, analog circuitry configured to implementEquation 2 may be used in place of the controller 68. The power driver60 contains the drive circuitry needed to convert the low level signalsfrom the controller 68 into the high power signals needed to drive themotor 56.

FIG. 5B depicts a practical mechanical configuration for implementingthe architecture shown in the FIG. 5A embodiment. In FIG. 5B, the torquesensor 1732 (which corresponds to ref. # 66 in FIG. 5A) is positioned atthe very top of the shank member 1716 (which corresponds to ref. # 52 inFIG. 5A).

Another approach for determining the ankle torque Γ_(A) is to break thattorque up into its constituent components, and analyze the torque ofeach of those components separately. For example, in the design depictedin FIG. 6A-C, there are two components that contribute to the totaltorque: the torque applied by the series elastic element (Γ_(S)) and thetorque applied by the bumper (Γ_(B)). The bumper is positioned betweenthe shank portion of the ankle and the foot portion, and can also beconsidered a hardstop when the stiffness is high. In alternativeembodiments, a spring may be used instead of a bumper. Note that theι_(B) component only comes into play during bumper engagement (i.e.,during dorsiflexion, when the shank member presses against a bumper thatis affixed to the foot member, or, in alternative embodiments, when thefoot member engages a bumper that is affixed to the shank member).

If each of the contributing components is known, the total ankle torquecan be determined by vector-adding Γ_(S) and Γ_(B) (i.e.,Γ_(A)=Γ_(S)+Γ_(B)). In the design depicted in FIG. 6B, both Γ_(S) andΓ_(B) can be determined as a function of displacement as measured byposition sensors that are distributed throughout the design, like amotor encoder that detects the position of the motor and an ankle angleencoder that detects the angle of the ankle pivot.

We begin with Γ_(S). In FIG. 6C, the motor 1B-102 drives a ballscrew1B-106, and a digital encoder 1B-104 mounted on the motor measures theballscrew extension p. If the foot were to be operated unloaded (e.g.,when it is up in the air), for every given value of ballscrew extensionp, the ankle joint 1B-108 would move to an angle β(p). The β(p) functioncan be determined empirically by lifting the device in the air so thatit is unloaded, then driving the motor through its entire operatingrange, and measuring the resulting angle of the ankle joint 1B-108 ateach value of p. Alternatively, β(p) could be calculated based on theknown geometry of the device. The β(p) function is stored in a memorythat is accessible by the controller 78 (shown in FIG. 6A) in anysuitable format (e.g., as an equation or a lookup table).

During normal operation, the device will be loaded, and the actual angleθ of the ankle joint 1B-108 can be determined (e.g., by ahigh-resolution encoder, not shown, mounted on the ankle joint). Inaddition, the actual ballscrew extension p can be determined based onthe output of the digital encoder 1B-104. The controller inputs p fromthe motor encoder and retrieves the unloaded angular position β(p) frommemory. It then inputs the actual angle θ from the ankle joint angleencoder and subtracts β(p) from 0 (i.e., the controller computesθ-β(p)). That difference is the angular deflection of the SEE 1B-110. Insome embodiments, a “single-turn” motor controller can be used. At poweron, its absolute position within one motor turn and the absolute jointposition can be used together to determine the absolute displacement ofthe ballscrew in relation to the end-of-travel in the plantarflexiondirection.

After the deflection has been determined, the torque Γ_(S) can be foundbecause torque is a function of the deflection. In a simple model, thetorque vs. deflection characteristics can be modeled as a linearfunction (Hooke's Law), so that Γ_(S)=k_(S)×deflection, where k_(S) isthe spring rate for the SEE. FIG. 9 depicts the torque vs. deflectioncharacteristics for the series elastic element 1B-110 (shown in FIG.6B). From these characteristics, a measured deflection can be used todetermine Γ_(S). Note that relying on an equation involving a springconstant k_(S) is just one of many possible ways to determine the torquefrom a deflection, and alternative models and approaches for determiningthe torque vs. deflection characteristics may also be used (e.g., alookup table, polynomial curve fitting, or non-linear estimation).

We turn next to the Γ_(B) component. During dorsiflexion, the shankmember 1B-111 pushes towards the foot member 1B-114, and a bumper 1B-112that sits between those two members (and could be affixed to eithermember) is compressed. During testing of the previous generationdesigns, which used a relatively soft plastic for the bumper 1B-112, theinventors recognized that there is observable compliance in the bumperduring engagement, in the range of 0.25° of deflection per 85 Nm peakreference load for a 250 lb amputee. When harder plastics are used(e.g., EPDM, with a 95 A durometer), there is much less deflection(e.g., 0.1° of deflection per 85 Nm peak reference load for a 250 lbamputee), and the force-deflection characteristic of this compliancebecame more stable and more easily modeled. Note that the metal shellsthat house the ankle mechanism will also flex measurably, and so can thefoot structure and the member that contacts the bumper. When theflexural displacements are measured empirically for a particular designor sample of a design (e.g., using a test fixture), all of thoseflexures would be automatically accounted for.

The variation of Γ_(B) with the compression of the bumper can bedetermined empirically for a given design or a particular instantiationof a design. One way to do this is to bolt a sample ankle/foot 250 intoa test fixture 200, like the one shown in FIG. 7. The test fixture 200preferably uses a six degree-of-freedom force-torque sensor 210 thatsimultaneously measures force and torque along and about threeorthogonal axes (e.g., made by JR3, Inc.), with a backdrive ballscrewactuator 220 installed between the foot portion 252 of the ankle/foot250 and the JR3 210. In this test fixture 200, the ankle/foot 250 isdriven until the foot portion 252 makes initial contact with the bumper(shown in FIG. 6B) on the shank portion 254 of the ankle/foot 250. Theangle of initial contact is defined as θ₁. Then, using the backdriveballscrew actuator 220, the foot portion 252 is further driven to anangle θ_(C). The angle θ_(C) can be measured by the ankle encoder 1B-108on the ankle/foot prosthesis (shown in FIG. 6C). As θ_(C) increases, thecompression of the bumper increases, and the forces as determined by theJR3 210 are stored for every possible angle θ_(C).

The Z (vertical) and Y (Horizontal) forces measured by the JR3 210 aresummed using vector mathematics to determine the force along thebackdrive screw axis. The ankle torque is then calculated by multiplyingthe axial force by the perpendicular moment arm, after subtracting anytorque contribution from the SEE. The ankle torque versus ankle angle isplotted for a number of cycles (e.g., 10 cycles) for every possibleangle θ_(C) and a least squares best fit line is calculated, assuming alinear relationship Γ_(B)=K_(S)×(θ_(C)−θ_(S)), where K_(S) is therotational spring rate for the bumper 1B-112. The slope of the resultingbest-fit line is the spring rate K_(S) of the bumper in Nm/rad as shownin FIG. 8A. In alternative embodiments, instead of using this linearrelationship to model the bumper, alternative models and approaches fordetermining the torque vs. deflection characteristics in the design mayalso be used (e.g., a lookup table, polynomial curve fitting, ornon-linear estimation).

Note that when increasing the torque (i.e., when the foot portion isbeing driven into the bumper and is compressing the bumper), therelationship of the ankle torque to ankle angle deflection is verylinear. However when returning back to zero (decreasing torque), thecurve is different. This discrepancy is due to the effect of the energyabsorbing properties of the bumper. It is preferable to use the slope ofthe least squares best fit line for the increasing torque portion todetermine the spring rate K_(S) of the bumper.

FIG. 8B depicts the Γ_(B) component of torque that is determined usingthis approach over time in a situation where the bumper is increasinglycompressed for about half a second (until the torque reaches −90), andthen released. The quantized nature of the Γ_(B) torque is a function ofthe encoder resolution. This quantization can be minimized by utilizinghigher resolution encoders. In one preferred embodiment, a 13 bitencoder (8196 counts/360 degrees) manufactured by Renishaw Inc (P/NRMB13BC1) is used. The Renishaw encoder employs a custom Hall-effect ICthat measures the field angle arising from a single-pole, cylindricalmagnet mounted on the foot structure in relation to the orientation ofthe IC affixed to a printed circuit assembly embedded in the ankleshell. Filtering of the angle measurement, using a FIR Low-Pass filterexecuting in a dedicated DSP, has been shown to extend the effectiveresolution to between 15-16 bits.

Once the torque vs. deflection characteristics of a bumper/ankle shellhas been modeled (e.g., as explained above), the Γ_(B) contribution atany given instant during operation of the prosthesis can be determinedby measuring θ_(C) and plugging the result into the equationΓ_(B)=K_(S)×(θ_(C)-θ_(I)), or into an alternative model that modelsΓ_(B) as a function of θ_(C). Thus, from a measured angular deflectionθ_(C), the second torque component Γ_(B) can be determined.

In alternative embodiments, other ankle angle encoding means could beemployed to determine how far the bumper has been compressed, includingoptical, magneto-restrictive and inductive sensors.

At this point, both the Γ_(S) and Γ_(B) components are known. Γ_(S) cannow be added to Γ_(B) to arrive at Γ_(A), and the resulting Γ_(A) isused as an input to Equation 2 to control the motor.

FIG. 6A is a system block diagram for implementing this approach bydetermining Γ_(S) and Γ_(B) separately and then adding those componentsto arrive at Γ_(A). Elements 52-60 are the same as the correspondinglynumbered elements in FIG. 5A. Angular position sensors 76 measure themotor displacement p and the ankle joint displacement θ, and sendoutputs representing those displacements to the controller 78. Thecontroller 78 is programmed to convert those displacements to torqueΓ_(S) as explained above. In addition, the controller 78 is programmedto convert the ankle joint displacement θ to torque Γ_(B) as explainedabove. The controller 78 then vector-adds Γ_(S) to Γ_(B) to determineΓ_(A). The controller 78 then controls the motor 56 (with the assistanceof the power driver 60, as in the FIG. 5A embodiment) by implementingEquation 2.

As mentioned above, n in Equation 2 can be tuned to make the device morecomfortable for the user. Other parameters may also be similarly tuned,such as pff and the threshold angular rate ω_(TH), which affects theKv(ω_(x)) function in Equation 2.

Referring now to FIG. 10, which is a Γ-Θ plot for the stance-phase,body-mass-normalized torque-angle, response of an intact ankle,additional parameters can be found that may be tuned in a prosthesis ororthosis to try to better mimic the intact ankle and thereby improvecomfort and performance. Examples include, modulating impedance as theankle-foot transitions from controlled plantar flexion (the slope ofK₁₋₂), through controlled dorsiflexion (the slope of K₂₋₃), to poweredplantarflexion (the slope of K₃₋₄). The initial values of these threeimpedances, and the initial value of θ at toe-off (θ*_(TOE-OFF)) can bederived from the mean Γ-Θ response of intact ankles, and those initialvalues can then be tuned to suit the activity level, limb length,body-mass distribution and preferences of an individual user.

In the above-described embodiments, a single motor is used to implementboth plantarflexion and dorsiflexion. But in alternative embodiments,that motor could be replaced by one motor for implementingplantarflexion, and another component for implementing dorsiflexion. Inother alternative embodiments, a plurality of motors may be arranged inparallel to perform both plantarflexion and dorsiflexion. In still otherembodiments, the electric motors described above can be replaced withother types of motors (e.g., hydraulic motors), in which case thecontroller and the power driver will have to be adjusted accordingly.

Note that while the concepts described above are explained in thecontext of prostheses, they can also be applied in the context oforthoses. In addition, while the embodiments described above all relateto ankles, the above-described concepts can be applied in otherprosthetic and orthotic applications, such as hips, torso, and arms, inwhich case suitable modification should be made that will be appreciatedby persons skilled in the relevant arts. For example, in the context ofa knee, where the reflex occurs right during toe-off, the walking speedprediction would use “fresh” shank speed measurement just prior totoe-off. In those other contexts, the shank member can be generalized asa proximal member, the foot member can be generalized as a distalmember, and dorsiflexion/plantarflexion can be generalized as varyingthe angle between the distal member and the proximal member. Theabove-described concepts can also be applied in the context of humanoidrobots.

While the present invention has been disclosed with reference to certainembodiments, numerous modifications, alterations, and changes to thedescribed embodiments are possible without departing from the sphere andscope of the present invention, as defined in the appended claims.Accordingly, it is intended that the present invention not be limited tothe described embodiments, but that it has the full scope defined by thelanguage of the following claims, and equivalents thereof.

What is claimed is:
 1. An ankle-foot prosthesis or orthosis apparatuscomprising: a shank member; a foot member that is operatively configuredwith respect to the shank member so as to support walking and permit thefoot member to plantarflex and dorsiflex with respect to the shankmember; a motor configured to plantarflex the foot member with respectto the shank member; a series elastic element connected between at leastone of (a) the motor and the shank member and (b) the motor and the footmember; at least one first sensor having an output from which a walkingspeed can be estimated; and a controller configured to control themotor's torque, based on the output of the at least one first sensor, sothat the motor's torque for slow walking speeds is lower than themotor's torque for fast walking speeds.
 2. The apparatus of claim 1,wherein the motor is also configured to dorsiflex the foot member withrespect to the shank member.
 3. The apparatus of claim 1, wherein the atleast one first sensor comprises at least one of an angular rate sensorand an IMU.
 4. The apparatus of claim 1, wherein the controller controlsthe motor's torque based on the output of the at least one first sensorimmediately before a reflex occurs.
 5. The apparatus of claim 1, whereinthe controller is configured to (i) determine, based on the output ofthe at least one first sensor, a control gain that varies with walkingspeed, wherein the control gain at slow walking speeds is lower than thecontrol gain at fast walking speeds, (ii) determine a desired motortorque based on the control gain and a determined ankle torque, and(iii) drive the motor to achieve the desired motor torque.
 6. Theapparatus of claim 1, wherein the controller is configured to (i)determine, based on the output of the at least one first sensor, anangular rate ω_(x) of the shank, (ii) determine a control gain Kv(ω_(x))that is a function of the angular rate, wherein the control gain at lowangular velocities is lower than the control gain at low angularvelocities, (iii) determine a desired motor torque based on the equationMotor torque=Kv(ω_(x))×pff×(normalized_Torque)^(n), where pff is aconstant and n is between 2 and 7, and (iv) drive the motor to achievethe desired motor torque.
 7. The apparatus of claim 6, whereinKv(ω_(x))=0 when ω_(x)=0, Kv(ω_(x))=1 when ω_(x) exceeds a thresholdω_(TH), and Kv(ω_(x)) is a monotonically increasing function betweenω_(x)=0 and ω_(TH).
 8. The apparatus of claim 7, wherein the motor isalso configured to dorsiflex the foot member with respect to the shankmember.
 9. The apparatus of claim 14, wherein the at least one secondsensor measures ankle torque directly.
 10. The apparatus of claim 14,wherein the at least one second sensor has at least one output fromwhich a deflection of series elastic element can be determined, and thecontroller computes the torque based on the at least one output.
 11. Theapparatus of claim 14, wherein the at least one second sensor comprisesa sensor that senses a position of the motor and a sensor that senses anangle of the foot member with respect to the shank member, and thecontroller computes the torque based on the sensed position of the motorand the sensed angle.
 12. The apparatus of claim 14, wherein the atleast one second sensor comprises a sensor that senses a position of themotor and a sensor that senses an angle of the foot member with respectto the shank member, and the controller determines a torque componentΓ_(S) based on the sensed position of the motor, the sensed angle, and atorque vs. deflection characteristics of the series elastic element. 13.The apparatus of claim 12, further comprising a bumper that iscompressed when the foot member is sufficiently dorsiflexed with respectto the shank member, wherein the controller determines a torquecomponent Γ_(B) based on the sensed angle and a torque vs. deflectioncharacteristics of the bumper, and wherein the controller determines atotal torque based on Γ_(S) and Γ_(B).
 14. The apparatus of claim 1,further comprising at least one second sensor having an output fromwhich ankle torque can be determined, and wherein the controller isconfigured to control the motor's torque based on the output of the atleast one second sensor.
 15. The apparatus of claim 1, wherein controlof the torque of the motor is based on non-linear positive feedbackaccording to the output of the at least one first sensor.
 16. Theapparatus of claim 1, wherein the non-linear positive feedback involvesadjustment of a parameter of a power control gain function, theparameter comprising at least one of a power control gain and a powerexponent.